Positron emission tomography (PET) is a medical imaging technique that is commonly used to produce three-dimensional images of subjects' bodies inner organs and parts. This technique hinges on detection of the two 511 keV γ-rays emitted as a consequence of annihilation of an electron present in the subjects' tissues with a positron emitted by radionuclides inoculated to subjects in form of specific tracers (e.g., 18F-loaded fluorodeoxyglucose 18F-FDG). The clinical analysis of interest determines the choice for the tracers, which are selected to target specific cells or proteins based as it is made possible by their affinity for the metabolic processes. Information resulting from the determination of the position of interactions of the two 511 keV γ-rays with the detector and, possibly, of their Time-Of-Flight (TOF), is eventually processed with tomographic methods by computer analysis to produce three-dimensional images of tracer distribution and concentrations on the bodies of the subjects, which are then evaluated for clinical purposes as proxies of distribution and concentrations of cancer cells or cells of other types.
The determination of the position of positron annihilation vertices is limited by the finite range of positrons of the two 511 keV γ-rays. For 18F, the range in tissues is 5 mm, but the Full Width at Half Maximum (FWHM) and the Full Width at Tenth of Maximum (FWTM) of the three-dimensional displacement from emission to annihilation points are contained within 0.1 mm and 0.5 mm respectively [1]. Another limiting factor is the acollinearity of the two 511 keV γ-rays resulting from the non-zero momentum of the positron-electron system undergoing annihilation. The angular distribution is approximately Gaussian with a FWHM of ˜0.5 degrees and limits the resolution in the axial and transaxial directions to Δx=0.5·D·tan (0.25°)=0.0022·D [1], where D is the ring diameter, of 80 cm for PET and TOF-PET units, corresponding to an intrinsic limit in the transaxial resolution of a few (˜2) mm [1].
Today's most sensitive commercial PET and TOF-PET (Time Of Flight Positron Emission Tomography) units for clinical studies achieve resolution for the 15 single event from positron annihilation of 4-5 mm in the axial and transaxial directions [2]. Today's commercial TOF-PET units however have a very poor resolution—a few cm—for the determination of the position of the single annihilation event in the radial direction, due to the sub-optimal performance of the TOF measurement for the 511 keV γ-rays. This sub-optimal performance of the TOF measurement, responsible for the very poor radial resolution, is ultimately due to the slow response of the inorganic scintillators used as positron detectors (e.g., LSO and LYSO with typical decay times in the range τ˜40-80 ns). The best figures achieved by commercial TOF-PET units for the TOF spread of the single annihilation events are in the range Δt˜500 ps, resulting in a radial resolution Δr=c Δt/2≃75 mm [1].
The poor resolution in the radial direction results in limitations to the resolution, contrast, and brightness of the clinical image. As detailed above, the TOF-PET units commercially available today cannot resolve the position of individual positron annihilation with resolution close to the two fundamental scales of lengths of the PET described above. For this reason, it is necessary to reconstruct clinical images by tomography: i.e., first determining the two-dimensional projections in surfaces perpendicular to the positrons' line of flight, which have an intrinsic precision of 4-5 mm (axial and transaxial directions), and then combining and fitting the two-dimensional projections obtained for different angles such as to obtain a three-dimensional image, whose ultimate resolution is typically of 4-5 mm.
The use of a tomographic procedure for the reconstruction of three-dimensional images is, by its own, a strong disadvantage of commercially available PET and TOF-PET units. Tomography requires combination of images with statistics of 18F decays much higher than would be otherwise required if the positron annihilation vertices were individually reconstructed with a linear resolution of the scale of the 18F positron range. As a consequence, the typical activity of 18F inoculated to patients is 10 mCi, responsible for a typical 12 mSv dose for each check-up procedure. While the dose is of limited statistical consequence for adult cancer patients, it is a major cause of concern in pediatric oncology, with young patients in remission potentially subjected to up to 20-30 check-up procedures in the course of their lives.
Another, additional, limitation of PET and TOF-PET units is their limited efficiency as γ-ray detectors and their limited energy resolution. The low efficiency in γ-ray detection stems from cost considerations: procurement of large surfaces of LSO and LYSO crystals is one of the driving costs, and this results in standard limitation of the thickness of crystals to no more than two interaction lengths. The limited energy resolution of 10-15% is due to the photoelectrons statistics at 511 keV energy afforded by inorganic crystals coupled with PhotoMultiplier Tubes (PMTs) or Silicon PhotoMultipliers (SiPMs). As discussed in the next paragraph, the limited γ-ray detection efficiency and the limited energy resolution both play a role in increasing the instrumental background.
A further limitation of PET and TOF-PET units is the low Signal-to-Noise Ratio (SNR). In PET jargon, events are subdivided in: 1) “T” events, defined as true coincidences of two 511 keV γ-rays originating from the same positron annihilation and fully adsorbed in the PET detector; 2) “S” events, defined as coincidences of 511 keV γ-rays originating from the same positron annihilation but releasing partial energy in a Compton scatter in the subjects' body, and with their cumulative energy only partially adsorbed in the PET detector; 3) “R” events, defined as random coincidence of two 511 keV γ-rays originating from two different positron annihilations. In PET and TOF-PET units, the SNR defined as the ratio T/(S+R) is in the range from 1:1 to 2:1. The limited energy resolution of PET and TOF-PET units inhibits the rejection of S events in the signal analysis phase. The limited γ-ray detection efficiency results in the necessity of usage of large 18F activities, in turn resulting in very high rate of R events (RR), typically much higher than the rate of T events (RT) at normal condition of operation for PET and TOF-PET exams.
PET and TOF-PET units can be combined with hadrotherapy machines for use as In-Beam Positron Emission Tomography (IB-PET) [3]. With the IB-PET technique, the PET or TOF-PET units are operated in combination with the beam line delivering protons or 12C nuclides for destruction of cancer cells. Detection of the 511 keV γ-rays produced by positron-emitting nuclides activated in subjects' the tissues directly by protons or 12C nuclides delivered to the subjects is used for monitoring of the dose delivered to patients. The known art allows measuring the overall dose administered to subjects with a coarse spatial approximation. However, to date no method has been devised to monitor online the dose delivered to cancer cells relative to that delivered to surrounding healthy tissues.
For all of the above reasons, there is a need for improving positron imaging techniques.